Positron Emission Tomography (PET) has gained significant popularity in nuclear medicine because of the ability to non-invasively study physiological processes within the body. PET is the most sensitive, and exhibits the greatest quantification accuracy, of any nuclear medicine imaging instrument available at the present time. Applications requiring this sensitivity and accuracy include those in the fields of oncology, cardiology and neurology.
Using compounds such as .sup.11 C-labeled glucose, .sup.18 F-labeled glucose, .sup.13 N-labeled ammonia and .sup.15 O-labeled water, PET can be used to study such physiological phenomena as blood flow, tissue viability, and in vivo brain neuron activity. Positrons emitted by these neutron deficient compounds interact with free electrons in the body area of interest, resulting in the annihilation of the positron. This annihilation yields the simultaneous emission of a pair of photons (gamma rays) approximately 180 degrees (angular) apart. A compound having the desired physiological effect is administered to the patient, and the radiation resulting from annihilation is detected by a PET tomograph. After acquiring these annihilation "event pairs" for a period of time, the isotope distribution in a cross section of the body can be reconstructed.
PET data acquisition occurs by detection of both photons emitted from the annihilation of the positron in a coincidence scheme. Due to the approximate 180 degree angle of departure from the annihilation site, the location of the two detectors registering the "event" define a chord passing through the location of the annihilation. By histogramming these lines of response (the chords), a "sinogram" is produced that may be used by a process of back-projection to produce a three dimensional image of the activity. Detection of these lines of activity is performed by a coincidence detection scheme. A valid event line is registered if both photons of an annihilation are detected within a coincidence window of time. Coincidence detection methods ensure (disregarding other second-order effects) that an event line is histogrammed only if both photons originate from the same positron annihilation.
In the traditional (2-D) acquisition of a modem PET tomograph, a collimator (usually tungsten) known as a septa is placed between the object within the field-of-view and the discrete axial rings of detectors. This septa limits the axial angle at which a gamma ray can impinge on a detector, typically limiting the number of axial rings of detectors that a given detector in a specific ring can form a coincidence with to a few rings toward the front of the tomograph from the given detector's ring, the same ring that the detector is within, and a few rings toward the rear of the tomograph from the given detector's ring. A more recent advancement in PET acquisition is 3-D, in which the septa are removed, which allows a given detector to be in coincidence with detectors from all other detector rings.
Another tomographic diagnostic system that is similar to PET is known as single photon emission computed tomography (SPECT). The distinction is that in SPECT, only a single photon from a nuclear decay within the patient is detected. Also, the line of response traveled by the photon is determined exclusively by detector collimation in SPECT, as opposed to the coincident detection of two collinear photons as in PET.
In computed axial tomography (CAT, or now also referred to as CT), an external x-ray source is caused to be passed around a patient. Detectors around the patient then respond to x-ray transmission through the patient to produce an image of an area of study. Unlike PET and SPECT, which are emission tomography techniques because they rely on detecting radiation emitted from the patient, CT is a transmission tomography technique which utilizes only a radiation source external to the patient.
The details of carrying out a PET study are given in numerous publications. Typically, the following references provide a background for PET. These are incorporated herein by reference for any of their teachings.
1. M. E. Phelps, et al.: "Positron Emission Tomography and Audiography", Raven Press, 1986;
2. R. D. Evans: "The Atomic Nucleus", Kreiger, 1955;
3. J. C. Moyers: "A High Performance Detector Electronics System for Positron Emission Tomography", Masters Thesis, University of Tennessee, Knoxville, Tenn., 1990;
4. U.S. Pat. No. 4,743,764 issued to M. E. Casey, et al, on May 10, 1988;
5. R. A. DeKemp, et al.: "Attenuation Correction in PET Using Single Photon Transmission Measurement", Med. Phys., vol. 21, 771-8, 1994;
6. S. R. Cherry, et al.: "3-D PET Using a Conventional Multislice Tomograph Without Septa", J1. C. A. T., 15(4) 655-668.
7. J. S. Karp, et al.: "Singles Transmission in Volume-Imaging PET With a .sup.137 Cs Source", Phys. Med. Biol. Vol. 40, 929-944 (1995).
8. S. K. Yu, et al.: "Single-Photon Transmission Measurements in Positron Tomography Using .sup.137 Cs", Phys. Med. Biol. Vol. 40, 1255-1266 (1995).
Both SPECT and CAT (or CT) systems are also well known to persons skilled in the art.
In order to achieve maximal quantitative measurement accuracy in tomography applications, an attenuation correction must be applied to the collected emission data. In a PET system, for example, this attenuation is dependent on both the total distance the two gamma rays must travel before striking the detector, and the density of the attenuating media in the path of travel. Depending on the location of the line of response within the patient's body, large variations in attenuating media cross section and density have to be traversed. If not corrected for, this attenuation causes unwanted spatial variations in the images that degrade the desired accuracy. As an example, for a cardiac study the attenuation is highest in the line of responses (LORs) passing through the width of the torso and arms, and attenuation is lowest in the LORs passing through from the front to the back of the chest.
Typically, the attenuation correction data in PET systems is produced by either: shape fitting and linear calculations using known attenuation constants, these being applicable to symmetric well-defined shapes such as the head and torso below the thorax (calculated attenuation); or through the measurement of the annihilation photon path's attenuation using a separate transmission scan (measured attenuation). The use of calculated attenuation correction, which introduces no statistical noise into the emission data, can be automated for simple geometries such as the head, and is the most prominent method used for brain studies. However, complexities in the attenuation media geometry within the chest have prevented the application of calculated attenuation from being practical for studies within this region of the body. Accordingly, transmission scanning has been utilized.
The total attenuation of a beam along a LOR through an object is equal to the attenuation that occurs for the two photons from an annihilation. Thus, the emission attenuation along the path can be measured by placing a source of gamma rays on the LOR outside of the body and measuring attenuation through the body along this line. It has been the practice to accomplish this attenuation measurement by placing a cylindrical positron emitter "sheet" within the PET tomograph's field of view (FOV) but outside of the region (the object) to be measured. The ratio of an already acquired blank scan (no object in the FOV) to the acquired transmission scan is calculated. These data represent the desired measured attenuation factors, which may vary spatially. These data are then applied to the emission data after a transmission scan of the object to correct for the spatial variations in attenuation.
There are two types of transmitter source units conventionally utilized in PET transmission scan data collection, both of which form a "sheet" of activity to surround the patient. One involves the placement of rings of activity aligned with detector rings around the inner face of the septa (see FIG. 1). The second type utilizes the rotation of one or more axially-oriented rods of activity in a circular path just inside the inner face of the septa (see FIG. 2).
The first of these two emitter systems (the ring source method) significantly reduces the sensitivity of the tomograph due to the close source-proximity dead time effects of the source activity on all of the detectors. Further, removal of this assembly is either performed manually by facility personnel or by a complex automated (more recent) mechanical assembly. Large, cumbersome, out of the FOV shielding is required for storage of the automated source when not in use, adding to the depth of the tomograph tunnel and, thus increasing incidence of patient claustrophobia. The second type of emitter, using rotating source(s) suffers from the above-mentioned problems and also, due to the shielding requirements, reduces the patient tunnel diameter, further increasing patient claustrophobia symptoms.
Both of the above automated source transportation methods suffer from high mechanical component cost and from low sensitivity. Due to the dead-time-induced reduction in tomograph sensitivity, lengthy acquisitions are required in order to achieve usable low noise transmission scan data.
Accordingly, it is an object of the present invention to provide a method and apparatus for rapidly moving a point source of radiation within a selected geometry to form attenuation correction data from radiation transmission measurements for correcting an emission data set, which may then be used to form an image within that geometry.
It is also an object of the present invention to provide a system that reduces the time of determining information in tomography scans.
It is another object of the present invention to provide an improved radiation emitter for carrying out attenuation data acquisition for use in obtaining increased accuracy in tomography scans.
Another object of the present invention is to provide for the controlling of a position of a point source of radiation and for determining that position so as to generate multi-dimensional attenuation correction data from radiation transmission.
It is still another object of the present invention to provide a radiation source of substantially increased activity that can be used in tomography applications.
A further object of the present invention is to provide an improved radiation emitter that requires no mechanical motion within a tomograph unit but accomplishes emission of radiation uniformly covering all detector coverage in cylindrical regions within the unit.
Another object of the present invention is to provide a method and apparatus for rapidly moving a point source of radiation within a selected geometry to form attenuation correction data from radiation transmission measurements for correcting an emission data set, wherein the point source of radiation is a CT scanner.
It is also an object of the present invention to provide a method and apparatus for using a single photon source having associated collimators to illuminate opposing, non-collimated, detectors.
These and other objects of the present invention will become apparent upon a consideration of the drawings forming a part of the disclosure of the invention, together with a complete description thereof that follows.